EP0749018B1

EP0749018B1 - A magnetic resonance imaging system

Info

Publication number: EP0749018B1
Application number: EP96303461A
Authority: EP
Inventors: Michael A. Morich; Labros S. Petropoulos
Original assignee: Marconi Medical Systems Inc
Current assignee: Philips Nuclear Medicine Inc
Priority date: 1995-06-12
Filing date: 1996-05-15
Publication date: 2001-09-12
Anticipated expiration: 2016-05-15
Also published as: DE69615092D1; DE69615092T2; US5585724A; EP0749018A1

Description

The present application relates to a magnetic resonance imaging system and a method of resonance imaging with this imaging system. It finds particular application in conjunction with magnetic resonance imaging of the torso of human subjects and will be described with particular reference thereto. It is to be appreciated, however, that the present invention also finds application in conjunction with imaging other parts of the human anatomy and with the imaging of non-human and inanimate subjects.
In magnetic resonance imaging, dipoles are selectively aligned with a primary magnetic field. Radio frequency excitation pulses are applied to stimulate resonance in the aligned dipoles and radio frequency magnetic resonance signals are collected from the resonating dipoles. Gradient magnetic field pulses are applied to encode spatial position. When imaging the upper torso, which includes the heart and other moving tissue, high speed image acquisition is advantageous.
To promote high speed image acquisition and high resolution, high strength magnetic field gradients with high slew rates are advantageous. That is, gradients of large magnitude that can be switched on and off very quickly improve data acquisition time and resolution. However, gradient strength varies inversely as the radius squared of the gradient coil and stored energy, a critical factor for slew rate, varies with the fifth power of the radius of the gradient coil.
To improve data acquisition speed at resolution in other parts of the human anatomy, smaller diameter gradient coils have been used, e.g., smaller diameter head or wrist coils. However, the width of the patient's shoulders has been a limiting consideration for upper torso imaging. To improve the magnetic field gradient characteristics in the upper torso, elliptical gradient coils and planar gradient coils have been utilized.
Typically, a head coil is on the order of 30 cm in diameter, whereas, a whole body gradient magnetic field coil is about 65 cm in diameter. Larger diameter gradient coils are typically positioned close to the main magnetic field coils in order to minimize main magnetic field coil diameters. When the gradient and main magnetic field coils are placed closely adjacent, the gradient coil assemblies tend to induce eddy currents in the main magnetic field coil assembly. Shield gradient coils are typically disposed between the gradient and main magnetic field coils to inhibit eddy currents. However, the shield gradient coils increase gradient coil power consumption, typically by about a factor of 2.
It has been proposed (US-A-5365927) for the gradient coil assembly of a magnetic resonance imaging system to include first and second coil portions physically separated by a gap, in the region of the head and lower torso of a patient, in order to allow access to a portion of the patient in an imaging volume, to assist a physician in medical procedures.
In accordance with the present invention, a magnetic resonance imaging system includes a primary magnet assembly for generating a temporally constant magnetic field through an examination region, a gradient coil assembly for inducing magnetic field gradients across the examination region, and a radio frequency coil for receiving resonance signals from the examination region, the gradient coil assembly including first and second gradient coil portions which are physically separated by a gap, the first and second coil portions being disposed on opposite sides of the examination region, the first and second gradient coil portions having windings for generating gradient magnetic fields along three mutually orthogonal axes, characterised in that at least one of the gradient coil portions has upper and lower sections which are selectively separable.
This facilitates patient receipt and exiting. The lower section of at least one gradient coil portion may be fixedly mounted to a patient support and the upper portion may be removable from the lower portion.
One of the gradient coil portions can extend around the patient's torso below the shoulders and the other can extend around the patient's head and neck above the shoulders. In this manner, an upper torso region of the patient can be imaged with a gradient coil assembly that is smaller in diameter than the width of the patient's shoulders. An invasive medical instrument can be mounted in a selected relationship to the subject for insertion through the interstitial gap. Preferably, the gradient coil assemblies are circularly cylindrical and the gap is at least 1/2 a diameter of the cylindrical gradient coil portions.
In accordance with another aspect of the present invention, a method of magnetic resonance imaging with a magnetic resonance imaging system includes a magnet assembly that defines an axial bore therein, a patient support which is selectively insertable into and removable from the axial bore, a gradient coil assembly having a first gradient coil assembly portion and a second gradient coil assembly portion which are separated by a gap, the first and second gradient coil assembly portions each having lower sections which are fixedly mounted to the patient support and upper sections which are selectively removable therefrom, and a radio frequency coil assembly having a first radio frequency coil assembly portion which is fixedly supported by the patient support between the first and second gradient coil assembly portions and an upper portion which is selectively positionable above an upper torso portion of the patient between the first and second gradient coil assembly portions, the method comprising: removing the gradient coil assembly portions upper sections; positioning the patient on the patient support with the patient's torso below the shoulders in one of the gradient coil assembly lower sections and with the patient's head and neck in the other of the gradient coil assembly lower sections; connecting the first and second gradient coil assembly upper sections to the first and second gradient assembly lower sections; disposing the radio frequency coil assembly upper portion over the patient's upper torso portion; positioning at least the upper torso portion of the patient and the first and second gradient coil assembly portions in the bore; conducting a magnetic resonance imaging sequence to generate an image representation of at least the upper torso portion of the patient disposed between the first and second gradient coil assembly portions; and selectively displaying at least a portion of the electronic image representation as a human-readable display.
One advantage of the present invention is that it improves the resolution of torso images.
Another advantage of the present invention resides in the improved imaging speeds and data acquisition times in torso imaging.
Another advantage of the present invention is that it facilitates access to the examined region of the patient, while the patient is still disposed in a known relationship to the gradient magnetic field coils.
The invention will now be further described, by way of example, with reference to the accompanying drawings in which:
FIGURE 1 is a diagrammatic illustration of a magnetic resonance imaging system in accordance with the present invention;
FIGURE 2 is an expanded view of the gradient and radio frequency coil assembly of FIGURE 1;
FIGURE 3 is a detailed view of one quadrant of an x or y-gradient coil distributed gradient coil winding;
FIGURE 4 illustrates a preferred distributed z-gradient coil winding pattern; and
FIGURE 5 is a plot of the stored energy versus the interstitial gap of gradient coil assemblies of the system in accordance with the present invention.
With reference to FIGURE 1, a magnetic resonance imaging system includes main or primary magnets 10. In the preferred embodiment, the main or primary magnets are annular superconducting magnets disposed adjacent opposite ends of the assembly within a liquid helium can 12. The liquid helium can and the magnets are surrounded by a plurality of cold shields 14 which assist in maintaining the superconducting magnets at superconducting temperatures while minimizing helium boil off. The magnet assembly is surrounded by a toroidal vacuum dewar 16. Between the magnets adjacent opposite sides of the center of the bore, windows 18 are cut through the vacuum dewar, cold shields, and helium can to provide access to the patient. The helium can, cold shields, and the vacuum dewar are, of course, closed around the windows to maintain their integrity.
The vacuum dewar and the magnet assembly define a central bore 20. A gradient coil assembly, preferably a self-shielded whole body gradient coil assembly 22 is disposed around an outer periphery of the bore 20. In the illustrated embodiment, the whole body gradient coil assembly 22 includes a primary gradient coil assembly 24 within the bore which includes x, y, and z-whole body gradient coil windings for generating magnetic field gradients along x, y, and z-directions. A shield gradient coil assembly 26 is disposed inside the vacuum dewar for cancelling the magnetic field gradients emanating toward the main magnet assembly. A radio frequency shield 30 lines an inner surface of the gradient coil assembly. The radio frequency shield is transparent to gradient (kHz) range magnetic fields but is opaque to radio frequency (MHz) signals. A whole body radio frequency coil 32, such as a birdcage style coil, is disposed around the inner surface of the radio frequency shield 30 surrounding the bore 20.
A retractable patient support 40 supports a subject to be examined, an insertable gradient coil assembly 42, and an insertable radio frequency coil assembly 44. The gradient coil assembly 42 includes a first or head gradient coil portion 42a and a second or lower abdominal cavity gradient coil portion 42b.
With continuing reference to FIGURE 1 and further reference to FIGURE 2, the first gradient coil portion 42a includes a lower portion 50a which is mounted to the patient support 40. An upper, selectively removable portion 52a is electrically and mechanically interconnected with the lower portion 50a for imaging and is released and removed to facilitate patient access. In the preferred embodiment, the upper and lower gradient coil assemblies are constructed on a rigid dielectric former. The upper and lower assemblies together carry a pair of x-gradient coil assemblies as illustrated in FIGURE 3 and a pair of y-gradient coil assemblies also as illustrated in FIGURE 3. The x and y-gradient coils are disposed 90° offset around a central axis of the bore 20. Z-gradient coils in the form of annular loops are also carried by the dielectric formers. In the preferred embodiment, the x and y-gradient coils are constructed of copper foil laminated to the dielectric former. Electrical connectors, such as metal pins and sockets are mounted in the upper and lower former portions for providing electrical continuity between coil portions on the upper and lower coil portions when the coil portions are assembled together.
The second gradient coil portion 42b also has a lower portion 50b mounted to the patient support 40 and an upper removable portion 52b. The upper and lower portions again are constructed of a rigid dielectric material to which x and y-gradient coils, preferably of the construction of FIGURE 3 are laminated along with annular z-gradient coils. Electrical pins or other connectors again provide electrical connection between coil portions on the upper and lower portions when assembled.
The radio frequency coil 44 in the illustrated embodiment, includes a lower saddle coil 60 mounted to the patient support closely adjacent and generally conforming to the patient's back. An upper saddle coil 62 is removably mounted closely adjacent to the patient's chest. The upper radio frequency coil preferably has an arc which generally conforms to the curvature of the patient's upper torso across the shoulders. Alternately, as illustrated in FIGURE 2, the radio frequency coil may be a circular coil with a diameter larger than the patient's shoulders, or an elliptical coil with its major diameter selected to receive a patient's shoulders and minor diameter dimensioned to receive the patient's chest.
In the preferred embodiment, the two gradient coil portions 42a, 42b each have a common diameter of 40-45 cm, with 42 cm being preferred. The coil portions 42a and 42b are separated from each other by up to 30 cm, with 20-30 cm of separation being preferred. In this manner, the separation between the gradient coil portions is preferably between 1/2 and 3/4 of the gradient coil diameter.
In operation, a sequence control 70 generates the appropriate gradient and radio frequency pulses of a selected magnetic resonance imaging sequence. More specifically, the characteristics of a selected magnetic resonance imaging sequence are withdrawn from a sequence memory 72 and stored and used to control a radio frequency sequence controller 74 and a gradient pulse controller 76. A common clock 78 clocks the radio frequency and gradient controllers simultaneously. The selected radio frequency pulse signals are conveyed to a radio frequency transmitter 80 which is selectively connectable to the whole body radio frequency coil 32 and the insertable radio frequency coil 44.
After magnetic resonance is induced, the insertable radio frequency coil 44 or a surface coil (not shown) receives the magnetic resonance signal and conveys it to a digital receiver 82. The digital receiver 82 demodulates and digitizes the magnetic resonance signal. A reconstruction processor 84 reconstructs the received magnetic resonance signals into a volumetric or slice image representation. A volume image memory 86 stores one or a series of image representations from the reconstruction processor. A video processor 88 converts selected portions of the image representations in the image memory 86 into appropriate format for display on a monitor 90. For example, the video processor may convert selected slices of an imaged volume into appropriate format for display. As another option, the video processor can select the corresponding slice in a series of temporally displaced images of the heart to provide a cine image representation which simulates a motion picture of the selected slice of the heart as the heart is beating. As yet another example, the video processor can assembly a three-dimensional rendering of a selected organ or region.
Based on the image representation, various diagnostic and medical procedure options are available. In the preferred embodiment, a non-ferrous medical instrument, such as a biopsy gun 100 is mounted to the patient support 40. The biopsy gun is adjustably positionable relative to the patient, e.g., its height, angle, or axial position. In the preferred embodiment, the radio frequency coil 44 is a saddle coil which is defined by conductors having large open regions therebetween. The medical instrument is positioned to pass through the open regions of the radio frequency coil. In one embodiment, electromechanical sensors sense the position adjustments of the biopsy gun to provide an electrical indication of the selected biopsy needle trajectory relative to a fixed coordinate system of the patient support 40. Magnetic resonance detectable elements, such as spheres of copper sulfate, are embedded in the patient support 40 under the central portion. These magnetic resonance markers produce characteristic, corresponding elements in the reconstructed volume image representations stored in the image memory 86. The magnetic resonance markers provide a fixed correlation between the reconstructed image representations and the preselected fixed frame of reference of the patient support 40. A frame of reference correlation processor 102 correlates the frame of reference of the image representations with the location of the biopsy gun 100 or other medical instrument. More specifically to the preferred embodiment, the correlation processor generates an indication of the trajectory of the biopsy gun relative to the frame of reference of the reconstructed video image and conveys such trajectory information to the image memory 86 and the video processor 88. The video processor superimposes the trajectory on the displayed image representation to provide a direct indication to the physician of the selected trajectory relative to the internal regions of the patient.
Various other techniques for correlating the position of the biopsy gun or other medical instrument with the frame of reference of the patient and the patient support 40 are also contemplated. For example, the medical instrument may carry sonic, optical, radio frequency, or other emitters which selectively emit signals that are received by corresponding receivers mounted to the patient support 40 in a known relationship. By triangulating the signals received by the receivers based on time of flight, angle of reception, and the like, the exact position of each emitter relative to the receivers, hence the coordinate system of the patient support 40, is readily calculable.
The window 18 enables the physician to position the biopsy gun or other medical instrument manually while monitoring its trajectory on the monitor 90. Alternately, the instrument may be positioned remotely using servomotors or the like. Once the desired trajectory is attained, other images through the trajectory are displayed to be sure that the trajectory does not pass through an inappropriate organ or tissue. Thereafter, the physician inserts a biopsy needle, for example, along the trajectory for a distance measured from the display on the monitor 90 to perform the biopsy. Analogously, other medical instruments may be inserted along the trajectory to perform other medical procedures. As yet another alternative, the biopsy gun may be spring loaded or mechanically driven. Other electrically controlled medical instruments may also be used.
In the preferred embodiment, the gradient coils generate a gradient magnetic field on the order of 80-100 mT/m with rise times on the order of 250 µsec. The generated magnetic field gradient should be linear in the region of interest centered between the coils, anti-symmetric around the origin along the z or longitudinal axis, and symmetric around the origin along x and y-axes orthogonal to the z-axis and each other. For a maximum gradient coil length of L and an interstitial gap of L₁, these gradient characteristics are achieved from a current distribution that is limited to the regions from -L/2 to -L₁/2 and from L₁/2 to L/2.
Looking first to the design of the z-gradient coil, the resultant magnetic field gradient varies linearly along the z-direction and is anti-symmetric around the origin at the geometric center of the cylinder. The z-component of the magnetic field is also homogeneous in the plane which is perpendicular to the z-axis, i.e., in the transverse direction. Because there is a linear relation between the current and the magnetic field, the current density of the coil satisfies similar symmetry conditions with the magnetic field. Further, because the gradient coil has a finite physical size, the current density is zero outside of the cylinder. The expression of the Fourier expansion around the geometric center for current for the gradient coil in this design is:
where j_n ^a are the Fourier coefficients in the expression and the sin k_Lnz cos k_1nz represents the anti-symmetry condition of the current around the origin. Prohibiting current in the gap L₁ and outside the-physical boundaries of the coil limits the values that k_Ln and k_1n can assume. Thus, the set of allowable values for k_Ln and k_1n are: j a (±L1 2 ) = 0 then k1n = (2n-1)πL1 j a (±L1 2 ) = 0 then kLn = 2nπL It is to be recognized by those of ordinary skill in the art that other end conditions can also be used, e.g., the current is not required to go to zero.
Because the expression of the magnetic field contains the Fourier transform of the current density, the Fourier transform of the current density which has the form of Equation (1) is calculated. The Fourier transform pairs for j_ ^a are defined as follows: j a(k) = ∫+∞ -∞j a(z)e-ikzdz j a (z) = 12π ∫+∞ -∞j a (k) e+ikzdk Considering the above-expressions, the Fourier transform of j_ ^a(z) is:
with:
It should be noted that the dependence of ψ_n(k) on the sign of k is: ψn(-k) = -ψn(k)
From the expression for the Fourier component of the current, the expression of the magnetic field in terms of the Fourier coefficients is:
Because ψ_n(k) is an odd function of k, the integral of Equation (7) is zero, unless the complex exponential is only an odd function of k. Accordingly, replacing e ^ikzby isin(kz), Equation (7) for the magnetic field can be expressed as:
Using the relationship:
the stored magnetic energy W_m is:
From the expressions of the magnetic field and the stored energy, the functional
is defined as:
Minimizing with respect to j_n ^a, a matrix equation for j_n' ^a is defined as:

Truncating the infinite summations at M terms, the matrix representation of Equation (12b) becomes: J a C = λ D or: J a = λ D C -1 where J ^a is a 1xM matrix or vector, C is an MxM matrix, λ is a 1xN matrix or vector, and D is a NxM matrix. The Lagrange multipliers can be found using the expression of the magnetic field. Then the matrix representation of the magnetic field is:
or: B z = J a D t where B _z is a 1xN matrix or vector, and the superscript t is the symbol for the transpose of the matrix. By replacing J ^a from Equation (13b) in Equation (14b), the expression of the magnetic field becomes: B z = λ D C -1 D t which leads to the determination of Lagrange multipliers as:
providing that the inverse matrix of the expression [DC ^-1 D ^t] exists. Upon determination of the Lagrange multipliers, the expression for the Fourier components of the current density in matrix form becomes:
In this manner, the expression for the Fourier components for the gradient coil is determined. Substituting back into the expression for current density, the continuous current distributions of the current density for the coil are readily determined.
The continuous current density is divided by a discretisation process with axially varying current densities to determine discrete current winding patterns for the coil. The continuous current distribution is divided into positive and negative current regions. Integrating the area underneath each region generates the total current which is contained in each region. Once the current for all regions of the cylinder is calculated, discrete current loops are placed in order to mimic the behavior of the continuous current pattern. Each of the regions is then filled with discrete wires carrying a prescribed amount of current which is the same for all wires. The most probable value of the current is found and the behavior of the screte current values is related closer to a continuous current density. More specifically, each region of the continuous current density is divided into smaller segments which correspond to equal amounts of current. Then, each wire or foil strip is placed through the middle of the segment in order to obtain an equal contribution from both sides of the segment. Alternately, one can calculate these values based on the center of mass technique described in U.S. Patent No. 5,296,810 of Morich.
Optionally, a shield gradient coil is provided to eliminate eddy current effects. In order to eliminate magnetic field gradients outside of the region of interest, while maintaining the above-described magnetic field conditions inside the region of interest, the relationship between the current density of the outer coil j_z ^b(k) and the inner coil j_z ^a(k) in the Fourier domain is: j b(k) = - aI1 (ka)bI1 (kb) j a(k) In this example, the kernels in the integral expressions for the total magnetic field and the stored magnetic energy are modified by multiplying with a factor S_ab defined as:
In the preferred embodiment, the gradient coil has a radius of 21 cm, a total length of 90 cm, and an interstitial gap of 30 cm. The Fourier series expansion was limited to 10 terms (M=10). Three constraint points were chosen to define the characteristics of the field inside a 25 cm spherical volume. The first constraint establishes a gradient field with a strength of 40 mT/m. The second constraint defines a linearity at a distance of 12.5 cm from the center of coil as varying not more than 5% of the ideal value, in the preferred embodiment. The third constraint defines the uniformity across the plane perpendicular to the gradient axis to be less than 15%, in the preferred embodiment. These constraints are displayed in TABLE 1.

n ρ_i z_i B_zSC(2n)

1 0.0000 0.001 0.00004000

2 0.0000 0.125 0.00500000

3 0.1250 0.001 0.00003440

Discretizing the current, a series of 18 negative and positive regions are defined as illustrated in FIGURE 4, which illustrates the axial spacing of winding loops which each carry a common current. The winding pattern illustrated in FIGURE 4 achieves a magnetic field gradient of 38.83 mT/m at 170 amps, a linearity along the gradient axis of 0.7% at z=±12.5 cm, and a uniformity of the gradient field inside the spherical volume of 13.57%. The calculated stored energy is 33.63 Joules at 100.5 mT/m gradient strength which achieves a rise time of t_r of 264.5 µsec. for a current of 440 Amps and a voltage of 600 volts. The corresponding slew rate is SR= 380 mT/m/msec. TABLE 2 illustrates a comparison of a conventional whole self-shielded body gradient coil (diameter =68 cm) and the above-described coil:

Property	CONVENTIONAL	INTERSTITIAL
Strength @ 440 Amps	28 mT/m	100.5 mT/m
SLEW RATE @ 600 Volts	75 mT/m/msec	380 mT/m/msec
Linearity @ z=±12.5 cm	0.7%	0.7%
Uniformity @ p=±12.5 cm	1.5%	13.6%

The transverse gradient coils again are designed by limiting the current to lie on the surface of the cylinder. The total current density for this current design can be represented as a superposition of two components, one along the axial and the other along the azimuthal direction. The current density is again limited to the region between the ends of the coil and the ends of the interstitial gap. For a transverse gradient coil, the total current density can be described as:
where δ(ρ-ρ_o) is the restriction that the current is defined on the cylindrical surface of radius ρ_o=a. The restriction on the total length of the coil, the confinement of the current density to the cylindrical surface, the azimuthal and axial symmetries for j_ ^a and j_z ^a and the requirement that the current density obeys the continuity equation provides a Fourier series expansion for both components around the geometric center of the coil as:

where j_n ^a are Fourier coefficients, L is the total length of the inner coil, and k_n= 2nπ/L, because the current cannot flow off the ends of the cylinder. The general expression of the magnetic field for a transverse gradient coil in terms of the Fourier transform of the current density is
where j_ ^a(m,k) is the double Fourier transform of j_ ^a(,z). Since the azimuthal dependence of j_ ^a is proportional to cos , the Fourier transform of j_ ^a is non-zero when m=±1. In this case, the two-dimensional Fourier transform of the current density is:
with:
where ψ_n(k) is an even function of k and j_ ^a(+1,k)=j_ ^a(-1,k). Thus, the expression of the gradient field has the form:
In a similar fashion, the stored magnetic energy is:
The restriction that the current density be confined to the interval of L₁/2≤|z| ≤L/2, the assumption that the function which represents the current density distribution is an analytical function of  and z, and the condition that the current density inside the interstitial gap is zero defines the condition for the z-component of the current density j _z ^a(,z) at the location ±L₁/2 as:
From the expressions of the magnetic field, the stored energy and the condition for the current density, the functional is constructed as:
where B_zSC and j_zSC ^a are the prespecified (constraint) values of the magnetic field at the constraint points and for the current density, respectively.
Minimizing with respect to j_n ^a, a matrix equation for j_n ^a is obtained as:
Truncating the infinite summation at M terms and using compact notation, Equation (30) reduces to:
or: J a C = λ D orJ A,= λ D C -1 where J ^a is a 1xM matrix or vector, C is a MxM matrix, λ is a 1xN₁+1 vector or matrix, and D is an N₁+1xM matrix with:
From Equation (32), the Fourier coefficients of the current density expansion can be determined to provide a continuous current distribution for the coil.
Optionally, to eliminate eddy currents, a second or shield gradient coil is provided to shield the region which is outside of both coils. In order for this condition to hold, the relationship between the current density of the outer coil j_ ^b(k) and the inner coil j_ ^a(k) in the Fourier domain is: j b (k) = - aI/ 1 (ka)bI/ 1 (kb) j a (k) The expressions for the total magnetic field and the stored magnetic energy are modified by multiplying the kernels of the various integrals with the factor S_ab which is defined as:
The discretization of the continuous current density can be implemented as follows. First, consider the continuity equation for the current density: ▿ ·J = 0 In analogy with the magnetic field, where a vector potential is introduced, the current density can be expressed as a curl of the function _s, called a "stream function". Specifically: J = ▿ × S Because the current density is restricted to flow on the surface of a cylinder of radius a=ρ_oand has only an angular and axial dependence, the relation between the current density and the stream function in cylindrical coordinates is: j a(, z) α  + jz a (, z) α z = ∂Sρ ∂z α  - 1a ∂Sρ ∂ α z and S_ρ is found from:
The contour plots of the current density are determined by:
where N is the number of current contours, S_min is the minimum value of the current density, and S_inc represents the amount of current between two contour lines. The determination of S_inc is: Sinc = Smax - Smin N with S_max representing the maximum value of the current density. The contours which are generated by this method follow the flow of current and the distance between them corresponds to a current equal to an amount of S_inc in Amps. Finally, discrete wires or foil strips are positioned in such a way as to coincide with these contour lines.
For a finite x-gradient coil with an interstitial gap, a cylinder radius of 21 cm, and a total length of 90 cm, and an interstitial gap of 20-30 cm, the gradient coil assumes the design as shown in FIGURE 3. More specifically, constraint points over a 25 cm spherical volume are: a gradient field strength of 40 mT/m linearity along the axis of the gradient of 15% and the third constraint sets the z-component of the current density at the boundaries of the interstitial gap to zero. The set of constraints shown in TABLE 3 is generated.

n ρ_i z_i B_zSC(2n)

1 0.0010 0.000 0.000040000

2 .125 0.0000 0.00425

3 j_zSC ^a = 0.000000001

Using these constraints in the aforementioned procedure, the Fourier coefficients are calculated. Hence, the continuous distribution of the current density for the x-gradient coil is obtained. Using the stream function technique, the discrete current distribution of the coil with the interstitial gap is obtained. The discrete version of the current pattern as illustrated in FIGURE 3 has 16 positive current loops which each carry about 258.5 Amps for 40 mT/m. The stored energy for this coil configuration is 8.36 Joules at the gradient strength of 40 mT/m. Assuming gradient amplifiers with true current output of 440 Amps. and a true voltage of 600 volts, the estimated rise time is t_r=206 µsec. at the gradient strength of 68.1 mT/m. The linearity inside a 25 cm spherical volume is about 15% and the uniformity of the gradient field inside the spherical volume is 20%. TABLE 4 displays a comparison between a conventional whole body self-shielded transverse gradient coil (diameter=68 cm) and the coil of FIGURE 3.

Property	CONVENTIONAL	INTERSTITIAL
Strength @ 440 Amps	27.3 mT/m	68.1 mT/m
SLEW RATE @ 600 Volts	73 mT/m/msec	331 mT/m/msec
Linearity @ p=±12.5 cm	2.5%	15%
Uniformity @ z=±12.5 cm	1.73%	20%

It will be appreciated that the gap may be selected to have various widths. As illustrated in FIGURE 5, the amount of stored energy varies in accordance with the length of the interstitial gap. It is to be appreciated that the stored energy with a gap of 20 cm is about 1/3 of the energy of a gap of 30 cm.
Various alternate embodiments immediately suggest themselves. For example, the gradient coil assembly may be elliptical rather than cylindrical. As yet another alternative, the gradient coil assembly may be planar or biplanar. Although distributed coils are preferred, the coils may also be bunched in nature. When the radio frequency coil is larger in diameter and placed outside of the gradient coils, a radio frequency shield may be positioned around the exterior of the gradient coil assembly.

Claims

A magnetic resonance imaging system which includes a primary magnet assembly (10) for generating a temporally constant magnetic field through an examination region, a gradient coil assembly (42) for inducing magnetic field gradients across the examination region, and a radio frequency coil (44) for receiving resonance signals from the examination region, the gradient coil assembly including first and second gradient coil portions (42a, 42b) which are physically separated by a gap, the first and second coil portions being disposed on opposite sides of the examination region, the first and second gradient coil portions (42a, 42b) having windings for generating gradient magnetic fields along three mutually orthogonal axes, characterised in that at least one of the gradient coil portions (42a, 42b) has upper (52a, 52b) and lower (50a, 50b) sections which are selectively separable.
A magnetic resonance imaging system as claimed in claim 1, wherein the first and second gradient coil portions (42a, 42b) are circularly cylindrical, the gap between the first and second gradient coil portions (42a, 42b) being at least 1/2 a diameter of the circularly cylindrical portions.
A magnetic resonance imaging system as claimed in claim 1 or claim 2, wherein: one of the gradient coil portions (42b) is configured to extend around the patient's torso below the shoulders; and the other gradient coil portion (42a) is configured to receive the patient's head.
A magnetic resonance imaging system as claimed in any one of the preceding claims, wherein the radio frequency coil (44) is disposed in the gap between the first and second gradient coil portions.
A magnetic resonance imaging system as claimed in any one of the preceding claims, including: a subject support (40) to which the first and second gradient coil portions (42a, 42b) are fixedly secured; and an invasive medical instrument (100) mounted in a selected relationship to a subject on the support (40) for insertion into the subject through the gap between the first and second gradient coil portions (42a, 42b).
A magnetic resonance imaging system as claimed in claim 5, further including means (102) for correlating coordinate systems of the invasive medical instrument (100) and the coordinate system of images generated by the magnetic resonance imaging system.
A magnetic resonance imaging system as claimed in any one of claims 1 to 6 further including a vacuum dewar (16) surrounding the primary magnet assembly, the vacuum dewar having at least one window (18) adjacent a central portion thereof for providing human access to a central portion of the examination region, whereby human access is provided through the window to a portion of the subject in the interstitial gap.
A method of magnetic resonance imaging with a magnetic resonance imaging system that includes a magnet assembly (10) that defines an axial bore (20) therein, a patient support (40) which is selectively insertable into and removable from the axial bore, a gradient coil assembly (42) having a first gradient coil assembly portion (42a) and a second gradient coil assembly portion (42b) which are separated by a gap, the first and second gradient coil assembly portions each having lower sections (50a, 50b) which are fixedly mounted to the patient support and upper sections (52a, 52b) which are selectively removable therefrom, and a radio frequency coil assembly (44) having a first radio frequency coil assembly portion (60) which is fixedly supported by the patient support (40) between the first and second gradient coil assembly portions and an upper portion (62) which is selectively positionable above an upper torso portion of the patient between the first and second gradient coil assembly portions (42a, 42b), the method comprising: removing the gradient coil assembly portions upper sections (52a, 52b); positioning the patient on the patient support (40) with the patient's torso below the shoulders in one of the gradient coil assembly lower sections and with the patient's head and neck in the other of the gradient coil assembly lower sections; connecting the first and second gradient coil assembly upper sections (52a, 52b) to the first and second gradient assembly lower sections (50a, 50b); disposing the radio frequency coil assembly upper portion (62) over the patient's upper torso portion; positioning at least the upper torso portion of the patient and the first and second gradient coil assembly portions (42a, 42b) in the bore; conducting a magnetic resonance imaging sequence to generate an image representation of at least the upper torso portion of the patient disposed between the first and second gradient coil assembly portions; and selectively displaying at least a portion of the electronic image representation as a human-readable display.